Biosensor technology has improved rapidly in the past decade because of its potential application in areas such as diagnosis of diseases, monitoring of clinical or environmental samples, fermentation and bio-processes, testing of pharmaceutical or food products, and early warning chemical and biological warfare detection systems. Generally, a biosensor is an analytical device, which uses biologically-sensitive materials for detecting biological or chemical species without the need for complex sample processing.
Biosensors are usually made by immobilizing a biologically sensitive material to a suitable transducer system, which converts the biochemical response into a quantifiable and processable electrical signal (Sethi R. S. Biosens. Bioelectron., 1994, 9, 243). The biologically-sensitive materials can be enzymes, multi-enzyme systems, organelles, membrane components, complete cells, antibodies or antigens, whole slices of mammalian or plant tissues, and the like. These materials are responsible for the recognition of the test or target species in the mixture that is analyzed and provide the selectivity and sensitivity of the final sensor device. For immobilizing these biological components onto the transducer surface, the transducer surface is first modified to form a suitable interface, e.g., by providing it with a suitable interface layer. This interface layer plays a key role in the biosensor performance. Several critical parameters of the biosensor, such as stability, reliability, sensitivity, and selectivity, ultimately depend on the properties of this interface layer (Byfield M. P. and Abuknesha in Biosens. Bioelectron.,1994, 9, 373).
For this purpose, self-assembled monolayers (SAMs) technology provides a powerful tool to generate an interface layer on a variety of substrates [Chaki N. K. and Vijayamohanan K. Biosens. Bioelectron. 2002, 17, 1]. For this kind of SAMs, a wide variety of chemical manipulation procedures can be utilized to achieve immobilization of the biological-sensitive materials in a quantitatively and spatially controllable way [Ostuni E. et al., Colloids and Surfaces B: Biointerfaces 1999, 15, 3], [Zammatteo N. et al., Anal. Biochem. 2000, 280, 143], [Shriver-Lake L. C., in: Immobilized Biomolecules in Analysis, Eds. Cass A. E. G. and Ligler F. A., Oxford University Press, Oxford, UK, 1998, 1-14].
For substrates based on, for example, gold, silver, platinum, copper and GaAs, alkane thiol SAMs are mostly used. However, there are several limitations to the use of these kinds of alkane thiol SAMs. Firstly, the stability of alkane thiol has always been an issue of debate [Gothelf K. V. and Larsen A. G., Journal of Colloid and Interface Science 2002, 255, 356], [Bearinger, J. P. et al., Nat. Mater. 2003, 2, 259]. Their limited robustness has limited their application in cell based biosensors (Mrksich M., Dike L. E., Tien J., Ingber D. E., Whitesides G. M. in Exp. Cell Res., 1997, 235, 305). Another important limitation is that the use of multi-component alkane thiol SAMs or mixed SAMs to control the density of the functional groups usually results in non-uniform structure caused by the micro-phase separation of different components [Brewer N. J. and Leggett G. J. Langmuir 2004, 20, 4109]. In addition, the density of the surface functional group cannot be easily controlled. It is assumed that the ratio of the two components in a mixed alkane thiol solution will determine the ratio of the two components in the resulting monolayer. However, this ratio may change when the two alkane thiol components, for example, chemisorb onto the substrate due to the possible adsorption kinetic difference of these two alkane thiols.
On the other hand, functional organo-silane SAMs are mainly used to modify oxide and glass surfaces. A main drawback of the organo-silane SAMs is that the density of the functional groups on the surface cannot be controlled easily. Attempts to control the functional group density either consist of complex surface modification procedures or result in non-uniform molecular structure [Buseman-Willams J. and Berg J. C., Langmuir 2004, 20, 2026].
Recently, a kind of self-assembled polymeric monolayer, which is able to chemisorb onto a gold surface through multiple thiol or disulfide groups grafted on the polymer backbone, has been reported. Compared to alkanethiol SAMs, this kind of self-assembled polymeric monolayer has improved stability by cooperative binding through multiple thiol-gold or disulfide-gold bonds [Johnson, P. A.; Levicky, R. Langmuir 2003, 19, 10288], [Sun, F. et al., J. Am. Chem. Soc. 1996, 118, 1856]. Levicky R. et al. demonstrated a thiol-derived polysiloxane, i.e. poly(mercaptopropyl)methylsiloxane, on the gold support. DNA oligonucleotides can be immobilized onto this polymeric interface with improved stability (Johnson, P. A.; Levicky, R., Langmuir 2003, 19, 10288). However, suitable cross-linker molecules had to be used to derive this polymeric interface in order to introduce suitable functional group onto the surfaces. The procedure complicates the surface modification and sometimes results in poor reproducibility. Meanwhile, the density of the surface functional groups cannot be controlled under these conditions.
Xia N. et al. reported an improved self-assembled polymeric monolayer on gold substrates (Xia, N.; Hu, Y. H.; Grainger, D. W.; Castner, D. G., Langmuir 2002, 18, 3255). A poly(methylhydrosiloxane) backbone with a molecular weight of 4500 was grafted with dialkyl disulfide chains, 600 Mw methoxy-terminated poly(ethylene glycol) (PEG) chains, and 3400 Mw PEG side chains terminated with N-hydroxysuccinimide (NHS) reactive ester groups. The ability of using this functionalized PEG-grafted polysiloxane monolayer to control protein binding was demonstrated. Despite the advantages reported by Xia et al., there are still several main drawbacks that could not be overcome easily. First, due to the limited solubility of this PEG-grafted polysiloxane polymer in other solvents, toluene has to be used as a solvent to prepare the polymeric monolayer. Toluene is a toxic solvent and thus disadvantageous to use when bio-related applications are considered. A lot of bio-molecules, such as short peptide (RGD for example), extra-cellular materials, biotin, and the like, which can serve as functional groups to immobilize the biologically-sensitive materials, cannot be integrated into this graft polysiloxane polymer due to the incompatibility of toluene with these bio-molecules. Using toluene as a solvent thus greatly limits the applications of this grafted polysiloxane polymer in the biosensor field. Furthermore, the architecture of the polymeric monolayer is not optimal when the polymeric monolayer is prepared from a toluene solution. For this PEG-grafted polysiloxane polymeric monolayer, the ideal architecture is a stratified structure, wherein the PEG side chains are enriched at the outermost surface to interact with biologically-sensitive material, the polysiloxane backbones present in between, and the alkyl disulfide chains stay close to the gold surface. However, when toluene is used as a solvent to prepare the polymeric monolayer, a lot of PEG side chains are buried under the polysiloxane backbones. At the same time, the surface coverage and packing density of the resulting polymeric monolayer are limited. These architectural problems are induced by the solvent effect on the organization of the PEG side chains and polysiloxane backbones. As a consequence, the nonspecific adsorption increases due to the exposed hydrophobic polysiloxane backbones and the defects in the polymeric monolayer, and the specific binding decreases because parts of the functional groups (active ester) are buried under the polysiloxane backbones.
In addition, for the polymer reported by Xia et al., there are two different kinds of PEG chains grafted onto the polysiloxane backbone. One kind of PEG chain has methoxy end groups, and is used to decrease the nonspecific adsorption. The other kind of PEG chain has active ester end groups, and is used to specifically immobilize a biomolecule. Counting the alkane disulfide chains also grafted onto the polysiloxane backbones, there is a total of three components that are grafted onto the polysiloxane backbones. For such a complex structure, the controllability of the graft ratio and the yield of the product are poor in the polymer synthesis step. Consequently, due to the poor controllability of the graft ratio, the surface density of the functional groups is not well controllable in the resulting polymeric monolayer. Moreover, due to the steric repulsion between the two kinds of PEG chains, the grafted ratio of the PEG chains with functional groups (active ester groups, for example) cannot go higher. This means that the surface density of the functional groups is limited to a low degree, which in turn significantly decreases the immobilization capacity of the surface. This is a big disadvantage in biosensor applications.